The ideal ultrasound system will maintain its imaging resolution at an optimum value throughout the area of interest. One method for accomplishing this is often referred to as beamformation with the complete data set or N.sup.2 reconstruction. With this method, the data acquisition sequence proceeds as follows: transmit with transducer element 1, receive with transducer elements 1 through N; transmit with transducer element 2, receive with transducer elements 1 through N; and so forth.
Since this approach requires N.sup.2 transmit/receive operations, it is not feasible for clinical imaging due to the data acquisition time requirements. However, it facilitates a beamformation process in which each individual pixel of the image has its own specific set of beamformation parameters. By this method, one can achieve dynamic focusing on transmit beamformation as well as with receive beamformation. Therefore, N.sup.2 reconstruction is often considered the target or the point of comparison by which clinically feasible approaches are measured. Thus a data acquisition method which approaches the N.sup.2 method while minimizing the frame rate impact would be a desirable development.
A conventional ultrasound image is composed of multiple image scan lines. A single scan line (or small localized group of scan lines) is acquired by transmitting ultrasound energy focused at a point in the region of interest, and receiving the reflected energy over time. The focused transmit energy is referred to as a transmit beam. During the time after transmit, one or more receive beamformers coherently sum the energy received by each channel, with dynamically changing phase rotation or delays, to produce peak sensitivity along the desired scan lines at ranges proportional to the elapsed time. The resulting focused sensitivity pattern is referred to as a receive beam. Resolution of each scan line is a result of the directivity of the associated transmit and receive beam pair.
The beamformer channel output signals are coherently summed to form a respective pixel intensity value for each sample volume in the object region or volume of interest. These pixel intensity values are log-compressed, scan-converted and then displayed as an image of the anatomy being scanned.
The frame rate of medical ultrasound imaging systems is determined by the number of transmit events necessary per frame. In conventional ultrasound imaging systems, a transmit event is a transmit beam directed in a particular direction or at a particular focal position. Frame rate in medical ultrasound imaging is a valuable resource. With additional frame rate, larger regions (as in color flow imaging or three-dimensional imaging) or faster objects (e.g., the heart) can be imaged. In addition, image enhancement methods such as video integration (noise reduction) or compounding (speckle reduction) can also use up frame rate.
In conventional medical ultrasound imaging, a single pulse is transmitted in a particular direction and the reflected echoes are coherently summed to form a single line in the image frame. The amount of time necessary to form that scan line is determined largely by the round-trip transit time of the ultrasonic pulse. Furthermore, many scan lines are present in an image frame to densely sample the anatomical region of interest. Thus the frame rate in conventional medical ultrasound imaging is determined by the sound propagation speed and the size of the region of interest.
High-frame-rate systems are desirable for 2D imaging and necessary for future real-time 3D imaging. The frame rate can be improved by decreasing the number of transmit events per frame. This has been conventionally accomplished with a proportional reduction in the number of transmit elements used in each transmit event, which results in poor signal-to-noise ratio (SNR). A decrease in the number of transmit events per frame has been conventionally accomplished only with an accompanying reduction in the number of transmit elements, which results in very poor SNR.
Conventional ultrasound beamformers use dynamic focusing during reception of echoes. With this method, the beamformation process is optimized for each depth to achieve as good a beamshape (i.e., narrow beamwidth with low sidelobes) as possible. However, in most systems, a single fixed focus is used during transmit beamformation to try to maintain a good combined beamshape. In areas away from the transmit focus, the beamwidth of the resultant beam widens and the sidelobes increase.
In one known ultrasound imaging system, an improvement to the focal properties is achieved by using multiple transmits aimed at different focal locations or zones. The echoes from these focal zones are used to form subimages, which then are stitched together in the final image. While this method optimizes beam properties in most areas of the image and hence begins to approximate N.sup.2 performance, this occurs at a major penalty of frame rate, i.e., the speed of sound is sufficiently slow to bring the frame rates down to as low as 5 frames/sec. In typical cases as many as eight transmit focal locations are used, which brings about an 8-fold reduction in frame rate. This penalty is quite severe with lower-frequency probes that are used in clinical situations requiring deep penetration.
A similar limitation associated with the data acquisition time occurs even more seriously with color flow mapping, a Doppler-based technique in which 4 to 16 transmissions are typically made in a direction of interest to acquire enough data for clinical utility. One approach that has been implemented to try to overcome this limitation is that of transmitting a wider beam and placing multiple receive beams within the transmit envelope. The resultant beams are not necessarily of good quality; however, given the relatively modest needs of Doppler processing, the method works satisfactorily. The quality of such beams is not sufficient for B-mode imaging. Hence, there is a significant need to acquire data at a faster rate and for that data to be of sufficient quality to form competitive images.